1. Field of the Invention
This invention relates broadly to optical systems. Particularly, this invention relates to micro-electro-mechanical-systems (MEMS) having optically reflective and transparent elements, and more particularly, to such systems wherein the optically transparent elements are transparent to ultraviolet light wavelengths.
2. State of the Art
The correction of abnormal human vision has progressed rapidly over the past few years. Although eyeglasses and contact lenses are still the dominant approach for correcting vision, newer techniques involving the reshaping of the cornea, and the replacement or supplement of the internal human lens, are providing more precise correction. To correct vision via the reshaping of the human cornea, precision surgical scalpels or lasers are used. Although still used, radial keratotomy (RK), which uses a surgical scalpel, is quickly being replaced by photorefractive keratectomy (PRK) and laser in-situ keratomileusis (LASIK), which use lasers. The laser refractive surgery field has exploded over the past few years with many new lasers and algorithms to correct human vision. Systems are now using laser wavelengths from the deep-ultraviolet to the infrared to change the shape of the cornea in a calculated pattern, which makes it possible for the eye to focus properly.
Artificial intraocular lenses (IOLs), which replace the human lens, usually due to cataract formation or lens damage, also provide very good vision correction, but, like contact lenses, which can offer “broad” spherical and astigmatic correction, they are limited in the precision of their corrective power. There are also supplemental refractive-IOLs which are inserted in the anterior chamber of the eye, the space between the iris and the cornea, while leaving the original lens intact. Light adjustable IOLs which are able to be altered outside the eye, or within the eye after implantation, thus allowing more custom-fit and more precise corrective power, are now undergoing research. Such an IOL is described in U.S. Pat. Application 2002/0016629 entitled “Application of Wavefront Sensor to Lenses Capable of Post-Fabrication Power Modification”, which is incorporated by reference herein in its entirety. Mid-UV to deep-UV lasers are used to alter these IOLs to provide varying refractive powers.
Finer, more precise measurements of human eye abnormalities have also been improving over the past several years. As these measurements have been improving, the industry has searched for ways to generate more custom corrections to the eye or to the IOL. The Digital Micromirror Device™ (DMD™), a micro-electro-mechanical-system (MEMS) semiconductor device consisting of hundreds of thousands of micromirrors, is an ideal device to deliver the custom laser beam pattern more precisely. The use of the DMD™ with respect to laser eye surgery is described in detail in co-owned U.S. Pat. No. 5,624,437, entitled “High Resolution, High Speed, Programmable Laser Beam Modulating Apparatus for Microsurgery”, U.S. Pat. No. 6,394,999 entitled “Laser Eye Surgery System Using Wavefront Sensor Analysis to Control Digital Micromirror (DMD) Mirror Patterns”, and U.S. Pat. No. 6,413,251 entitled “Method and System for Controlling a Digital Micromirror Device for Laser Refractive Eye Surgery”, all of which are hereby incorporated by reference herein in their entireties. Current, commercially-available DMD devices are designed to deliver visible wavelengths (from 400-nm to 750-nm), however, and cannot be used to deliver a large range of UV energy because of their protective window, which environmentally guards the micromirrors. UV energy can be categorized by wavelength according to physical definitions: extreme UV (EUV)(10 nm to 100 nm), vacuum UV (VUV)(10 nm to 200 nm, with recognition that VUV overlaps EUV), far or deep UV (DUV)(200 nm to 300 nm), and near UV (NUV)(300 nm to 400 nm). In addition, UV energy can be categorized by wavelength according to photobiologic definitions: UV-C (100 nm to 280 nm) which overlaps far and deep UV, UV-B (280 nm to 315 nm) which overlaps far and near UV and is also termed mid-UV, and UV-A (315 nm to 400 nm) which overlaps deep and near UV and which is also termed near-UV for photobiologic purposes.
Refractive Surgery: Corneal Reshaping by Laser and the Use of the DMD
Initial systems approved by the FDA for corneal reshaping implement the refractive correction by delivering beam-shaped laser energy based on first-order approximations of refraction from a single spherical surface. These systems implement a “broadbeam” approach, whereby the laser beam is shaped by a motorized iris (myopia and hyperopia) and motorized slit (astigmatism) based on profiles derived through Munnerlyn's derivation, as discussed in C. R. Munnerlyn, et al., “Photorefractive keratectomy: a technique for laser refractive surgery,” J. Cataract Refract. Surg. 14, 46–52 (1988). Typical systems using this approach on the market are VISX and Summit. More than one million eyes have been treated in this manner. This system is limited, however, as it treats a broad area of the cornea all at one time. Eye topography maps and more recently, wavefront analysis, reveal the eye has many minute variations across the cornea. See, e.g., J. Liang, et al., “Objective measurement of wave aberrations of the human eye with the use of a Hartmann-Shack wave-front sensor,” J. Opt. Soc. Am. A, Vol. 11, No. 7, 1949–1957 (1994). Referring to FIGS. 1(a) and 1(b), the broadbeam laser approach cannot correct these minute variations.
The latest systems being introduced to the market are based on scanning a small laser spot (typically 0.5-mm to 1.0-mm diameter), or a combination of different sized spots, across the cornea in a predetermined pattern to achieve refractive corrections, and are termed “scanning spot systems”. These scanning spot systems differ in that they are more flexible than the broadbeam approach. Referring to FIG. 2, with the control of a small spot 10, one can shape different areas of the cornea 12 independently of other areas. These techniques allow for a more general pattern to be applied to the cornea. Typical systems using the scanning spot approach are VISX, Autonomous Technologies, LaserSight, and SurgiLight.
However, there are several problems with the scanning spot approach when compared with the broadbeam approach but the flexibility offered appears to outweigh these. Some problems include longer refractive surgery time (speed), safety, tracking and surface roughness, which are discussed in more detail as follows.
With respect to longer refractive surgery times, the scanning spot is a slower approach since a small spot (typically 1-mm diameter) has to be moved over a wide surface (up to 10-mm for hyperopia). The broadbeam approach treats the entire cornea for each laser pulse, or treatment slice. The scanning spot system must deliver several hundred spots per treatment slice; thus, treatment times can increase.
With respect to safety, the broadbeam laser is inherently safe from a treatment interruption standpoint because the cornea is treated symmetrically for each pulse (the iris represents a circle and the slit represents a rectangle so that every point on the cornea is treated the same for each laser pulse). If the procedure is interrupted, you are guaranteed to have some symmetrical spherical or cylindrical correction, which can be continued easily. The scanning spot, with its small spot size, cannot cover the entire corneal surface with one laser pulse so that if an interruption occurs, there is no guarantee of a symmetrical etch at that point.
With respect to tracking, in the scanning spot system, the eye needs to be tracked in order to deliver the spot to the correct point on the cornea as the eye moves. This is not as much of a problem in the broadbeam system as a broader area is treated with each pulse.
With respect to surface roughness, laser spot overlap tends to create roughness in the resulting etch. While it is necessary to overlap spots to provide complete coverage for a given ablation zone, regions of overlap will be ablated at twice the etch depth per pulse. The smoothness of the ablated volume is dependent on the spot overlap and to a lesser extent, the ratio of spot diameter and ablation zone diameter. This problem is not seen in the broadbeam approach.
More recently, eye contour topography is being used to more accurately provide refractive measurements. Current FDA-approved refractive laser systems do not directly use eye-modeling systems, such as corneal topographers or wavefront sensors, to create the correct treatment profile for the patient's eye. The topographic map is used indirectly by the surgeon for optimizing the treatment plan (diopter correction and astigmatic axis). There are systems currently going through FDA trials that do use corneal topographic surface data to directly guide the laser treatment algorithm. This is accomplished using some type of scanning laser spot, as current broadbeam laser refractive surgery systems cannot provide the laser beam detail required to use the topographic map data. Each eye is individually analyzed as to its contour before ablation is applied. The idea here is to take into account the varying degrees of curvature and height variations across the corneal surface, as opposed to assuming a spherical surface as is currently done in broadbeam systems. Once these curvatures, or powers, are determined by eye topography, e.g., by using a system sold by Keratron, Orbtek, or Zeiss-Humphrey, they can be considered within the refraction correction derivation (as described by Munnerlyn) to create a customized ablation pattern for each individual eye. These ablations must be implemented by a scanning spot system, or better yet a DMD approach, as individual areas must be treated differently than other areas. Previously incorporated U.S. Pat. Nos. 5,624,437 and 6,413,251 discuss the DMD approach. Even this approach is limited in that only aberrations measured on the corneal surface are included in the refractive correction derivation.
The optimum approach to date uses the recent introduction of wavefront sensing analysis of the eye. With this new technology, a very powerful set of tools to correct the eye corneal surface is provided. Wavefront sensing provides an overall refractive analysis of the entire eye optical system, e.g., taking into account the cornea, the lens, the vitreous and the retina. The result of a wavefront sensor analysis yields a waveform model that represents a nearly perfect refraction measurement. This provides a superior analysis of the eye versus the current topography systems that only analyze the cornea. Wavefront data may be used to drive a scanning spot system, but it still encompasses the problems discussed before. However, it is directly compatible with the DMD approach due to the digital nature of the wavefront sensor analysis. This wavefront sensor-DMD approach is discussed in detail in previously incorporated U.S. Pat. No. 6,394,999.
Refractive Surgery: Light Activated IOLs and the Use of the DMD
Modern cataract surgery, in which the cataract is actually extracted from the eye, was introduced by Jacques Daviel in Paris in 1748. Samuel Sharp of London introduced the concept of intracapsular cataract surgery in 1753 by using pressure with his thumb to remove the entire lens intact through an incision. Small suction cups (erysiphakes) were introduced for this purpose in 1902 as well as various capsular forceps to grasp the lens for removal. Henry Willard Williams of Boston first described the use of sutures for cataract surgery in 1867. In the 1840s general anesthesia was introduced for surgical procedures, however it was not until 1884 that anesthesia in the form of eye drops (cocaine) was developed, obviating the hazards of general anesthesia and its postoperative complications. After Harold Ridley introduced the intraocular lens in England in the 1940s, efficient and comfortable visual rehabilitation became possible following cataract surgery.
The intraocular lens, or IOL, is a permanent plastic lens implanted inside the eye to replace the crystalline lens. In 1957 Barraquer of Spain used alpha-chymotrypsin to enzymatically dissolve the zonules for removal of the lens. Cryosurgery was introduced by Krawicz of Poland in 1961 to remove the lens with a tiny probe that could attach by freezing a small area on the surface of the cataract. In the late 1960s Charles Kelman of New York developed a technique for emulsifying the lens contents using ultrasonic vibrations and aspirating the emulsified cataract. In recent decades, there has been a rapid evolution of designs, materials, and implantation techniques for intraocular lenses, making them a safe and practical way to restore normal vision at the time of surgery.
More recently, designs have been implemented to provide accommodation with an IOL. Such attempts include diffractive multifocal lenses, flexible (fluid-filled) lenses, multi-element designs and hinged optics. These IOLs still only offer “broad” fixed power correction (both spherical and astigmatic), although they do offer some accommodative power.
Even more recently, light-activated IOLs have been proposed. These optical elements have a refraction modulating composition dispersed in a polymer matrix. The refractive modulating composition is capable of stimulus-induced polymerization, e.g., a UV light stimulus (longer wavelength UV: 325-nm to 340-nm range). In this way, an optical measurement (e.g., topography, wavefront, etc.) of the eye can be made followed by inducing an amount of polymerization of the refractive modulating composition, wherein the amount of polymerization is determined by the optical measurement. Thus, the varying degrees of curvature and height variations across the corneal surface can be taken into account, as opposed to assuming a simple spherical surface. Currently, collimated light from a Xe:Hg arc lamp (340-nm through a 1-mm photomask), or collimated light from a He:Cd laser (325-nm, 1-mm beam diameter), is used to activate and stabilize (or “lock in”) the refractive modulating composition. To polymerize the entire IOL, the 1-mm photomask, or 1-mm laser beam, must be moved or scanned across the IOL, as described in U.S. Pat. Application 2002/0016629 A1 entitled “Application of Wavefront Sensor to Lenses Capable of Post-Fabrication Power Modification”, which is incorporated by reference herein in its entirety. Scanning the small masks or small diameter laser beams across the IOL present the same problems as described in the scanning spot approach to corneal tissue reshaping discussed above. Therefore, when coupled to a proper collimated broadbeam arc lamp, or a broadbeam laser, this technique is directly compatible with the DMD approach due to the digital nature of the topographic or wavefront sensor analysis, and since it can provide larger, custom laser beam patterns more precisely.
The DMD Problem for Ultraviolet Wavelengths
From the above discussion, it is apparent that the DMD is ideally suited to provide the necessary laser delivery control, in both the corneal reshaping application and the IOL activation application, to obtain the finer resolution and custom laser beam patterns required by the more advanced measurement techniques. To date, however, there has not been a commercially-available DMD to provide delivery of the excimer wavelength used in laser refractive surgery (193-nm) or the wavelengths used in the light-activated IOL approach (325-nm to 340-nm). The shortest optimized wavelength able to be delivered by a manufactured, but still experimental DMD is 365-nm, far above the 193-nm necessary to etch the cornea in the laser refractive surgery area.
Referring to FIG. 3, it is noted that at 325-nm, a wavelength used in the IOL application, only about 83% of the light is transmitted through the UV-coated DMD window in a single pass. Moreover, with the DMD device, the light that strikes the mirrors must make two passes through the window; that is, the light must travel through the window first, reflect from the mirrors and travel through the window again to exit the device. This means that only about 69% (0.83×0.83) of the original 325-nm light returns from the DMD, assuming 100% reflection from the DMD mirrors, or a loss of 31% due to the window alone. At 340-nm, only about 91% of the light is transmitted in a single pass. Thus, for 340-nm, only about 83% (0.91×0.91) of the original 340-nm light returns from the DMD, assuming 100% reflection from the DMD mirrors, or a loss of 17% due to the window alone. The result of this loss is that the laser source must operate at higher energy output levels, which also increases damage to the beam shaping and delivery optics. Furthermore, there are other losses associated with the optics that are required to shape, homogenize, and deliver the laser beam (typically 45% to 60% loss). Additional losses result from damage to the optical coatings as a result of use.
To further illustrate the window losses, consider an example from U.S. Pat. Application 2002/0016629 A1, where a 1-mm diameter, 325-nm He—Cd laser beam with an energy density of 257 mJ/cm2 is required to induce the refractive modulating composition polymerization. To cover the entire 6.35-mm IOL with the laser beam at one time, as would be necessary using a DMD to activate the refractive modulating composition, this would require 81 mJ of energy at the IOL. With the above window losses of 31% at 325-nm, coupled to the typical losses of 50% for the beam shaping and delivery optics, this would require over 500-mJ from the laser, while operating from two minutes to ten minutes. Thus, it would be advantageous to have a deep UV window as lossless as possible to keep the laser energy requirement down.
As another example, consider the energy density required to etch corneal tissue. Although energy densities vary from system to system, a typical value is 160 mJ/cm2. A current commercially-manufactured, but experimental, DMD will not work for the laser refractive surgery wavelengths of between 190-nm to 250-nm (typically 193-nm) because as seen from FIG. 3, the UV-coated window covering the mirrors of the DMD has a 0% optical transmission below 250-nm. Thus, the only way to implement the DMD for this application is to use a window designed for these deep-UV wavelengths. For a typical 6-mm spot used in laser refractive surgery, an energy density of 160 mJ/cm2 requires 45-mJ. For a 10-mm spot, used to correct hyperopia, 125-mJ is required. Typical treatment times for broadband lasers range from a few seconds to one minute. Thus, the deep UV window needs to be as lossless as possible to keep the laser energy requirement down.
Another significant problem that exists with current DMD window designs is the reflections from the window surfaces, particularly on the DMD mirror side of the window. If the window and its coating are not optimized for the wavelength used, reflections can bounce back and forth from the inside of the window to the mirrors giving rise to “ghost images” that can cause interference problems resulting in incorrect patterns being delivered.
Current DMD windows assemblies are constructed of a Kovar® (ASTM-F-15) metal alloy frame and a borosilicate glass window. This combination is a common glass-ceramic sealing systems for protecting semiconductors (e.g., EPROMS) from a local environment. Kovar® is a low-expansion alloy whose chemical composition is controlled within narrow limits to assure precise uniform thermal expansion properties.
The most common borosilicate glass used in the current DMD application is Corning 7056. Corning 7056 glass works well for the visible light spectrum DMD application because it passes visible light well and its coefficient of thermal expansion (CTE) is very close to Kovar® (Corning 7056: 5.15×10−6/° C. versus Kovar®: 5.2×10−6/° C.). This allows a glass-to-metal hermetic seal when both are heated to nearly 1000 degrees Celsius. The traditional hermeticity definition is based on the Helium Fine Leak Test (mil-std 803 or JEDIC-JESD22-A109-A) where the value must be 5×10−8 atm-cc/s helium or better. The hermetic seal is formed by heating both the glass and metal until a wetting of the metal by the glass occurs, followed by the development of a chemical bond or some mechanical interlocking, thus maintaining the seal. The base transmission spectrum of Corning 7056 is shown in FIG. 4. By applying appropriate anti-reflection (AR) coatings to the glass, the transmission spectrums of FIG. 3 are achieved. Note the optical transmission can be shifted lower to handle the near-UV wavelengths, but this degrades part of the visible spectrum. Finally, note again that this glass will not pass deep-UV to mid-UV wavelengths very effectively. Instead a different material, such as fused silica, is required.
Fused silica is one of the most common materials used in the deep-UV to mid-UV applications. Fused silica is a polycrystalline, isotropic material with no crystal orientation. Its physical, thermal, dielectric and optical properties are uniform in all directions of measurement. There are special grades of fused silica, termed excimer grade fused silica, made especially for the above applications. Unfortunately, the CTE of fused silica (0.55×10−6/° C.) is not very close to the CTE of Kovar® (5.2×10−6/° C.) (differing by substantially an order of magnitude), and thus during the manufacturing process, and in the post-manufacturing environment, as temperatures vary, the hermetic seal between the two is not maintained. This allows the outside environment into the hermetically-sealed DMD semiconductor space and this becomes detrimental to the micromirrors behind the window causing them not to function properly. One of the most common problems is that the mirrors stick and do not respond to commands.
There are three additional problems with respect to current DMD mirror design. First, the current commercially-available DMD device uses bare aluminum mirrors to reflect the incoming light. Uncoated, or bare, aluminum provides about an 85% absolute reflectance from 200-nm to 2000-nm. This reflectance increases as wavelengths move into the visible area (about 90% averaged over 400-nm to 750-nm), as in the main application of the currently-manufactured DMD devices, and decreases as wavelengths move below 200-nm (e.g., about 84–86% for 193-nm and about 65–70% for 157-nm). Therefore, the uncoated aluminum DMD mirrors provide less than 85% reflectance, or greater than a 15% loss, for certain UV applications.
Second, the UV energy that strikes the mirrors also gradually erodes the mirrors, which damages or deforms them. This negatively affects the laser energy pattern that is delivered to the target.
Third, the incoming UV energy will travel between the mirrors and impinges on the underlying semiconductor control structure behind the mirrors. This can lead to degradation of the DMD device.